Field of the Invention
The present invention concerns a method for the simultaneous generation of scan data representing different contrasts of an examination object by magnetic resonance technology.
Description of the Prior Art
Magnetic resonance (MR) technology is a known technology with which images of the interior of an examination object can be generated. Expressed simply, the examination object is positioned in a magnetic resonance scanner in a strong, static, homogeneous basic magnetic field, also known as the B0 field, with field strengths of 0.2 tesla to 7 tesla or more, so that nuclear spins orient in the object along the basic magnetic field. In order to trigger nuclear resonances, radio-frequency excitation pulses (RF pulses) are radiated into the examination object and the nuclear spin resonances triggered thereby are detected as so-called k-space data and, on the basis thereof, MR images are reconstructed or spectroscopic data is determined. For spatially encoding the scan data, rapidly switched magnetic gradient fields are overlaid on the basic magnetic field. The scan data recorded are digitized and stored as complex numerical values in a k-space matrix. From the k-space matrix populated with such values, an associated MR image is reconstructed, such as by a multi-dimensional Fourier transform.
The desire for ever faster MR recordings in the clinical environment is currently leading to a resurgence of methods in which a number of images are recorded simultaneously. In general, these methods can be characterized by at least during a part of the scan, targeted transverse magnetization of at least two slices is used simultaneously for the imaging process (“multi-slice imaging” or “slice multiplexing”). In contrast thereto, in established multi-slice imaging, the signal is recorded from at least two slices alternatingly, i.e. completely independently of one another with correspondingly long scan times.
Known methods for such imaging are, for example, Hadamard encoding, methods with simultaneous echo refocusing, methods with broadband data recording, and methods that use parallel imaging in the slice direction. The latter methods include, for example, the CAIPIRINHA technique, as described by Breuer et al. in “Controlled Aliasing in Parallel Imaging Results in Higher Acceleration (CAIPIRINHA) for Multi-Slice Imaging”, Magnetic Resonance in Medicine 53, 2005, pp. 684-691, and the blipped CAIPIRINHA technique as described by Setsompop et al. in “Blipped-Controlled Aliasing in Parallel Imaging for Simultaneous Multislice Echo Planar Imaging With Reduced g-Factor Penalty”, Magnetic Resonance in Medicine 67, 2012, pp. 1210-1224.
Particularly in the latter slice multiplexing method, a so-called multi-band RF pulse is used in order to excite two or more slices simultaneously or otherwise manipulate them, e.g. to refocus or saturate them. Such a multi-band RF pulse is, a multiplex of individual RF pulses, which would otherwise be used for manipulation of the individual slices to be manipulated simultaneously. By such multiplexing, a baseband-modulated multi-band RF pulse is obtained from the sum of the pulse forms of the individual RF pulses. The spatial encoding of the recorded signals in two directions (two-dimensional gradient encoding) is achieved by a commonly used gradient circuit. It is, however, also possible to excite and manipulate each of the slices to be excited from which echo signals are simultaneously recorded, i.e. within a recording of scan data, with individual, e.g. successively switched, RF pulses. A combined use of “individual slice” RF pulses and multi-band RF pulses is also possible.
The signals arising from all the excited slices are recorded collapsed in one data set by multiple receiving antennae, and are then separated according to the individual slices with the use of parallel acquisition techniques.
The aforementioned parallel acquisition techniques (PPA techniques), with which in general acquisition times for recording the desired data can be shortened by sampling that is incomplete according to Nyquist, i.e. an underscan (undersampling), of k-space include, for example, GRAPPA (“GeneRalized Autocalibrating Partially Parallel Acquisition”) and SENSE (“SENSitivity Encoding”). In parallel acquisition techniques, the scan points in k-space that are not scanned (filled with acquired data values) during the undersampling are typically evenly distributed over k-space to be scanned according to Nyquist, so that, for example, every second (acceleration factor 2) or third (acceleration factor 3), etc., k-space row is scanned. In addition, the “missing” k-space data are reconstructed in parallel acquisition techniques with the use of coil sensitivity data. This coil sensitivity data of the receiving coils used during the recording of the scan data are determined from reference scan data, which samples fully according to Nyquist at least a region of k-space to be scanned, typically the central region.
In slice multiplexing methods, parallel acquisition techniques are used in order to separate again the scan data recorded simultaneously for different slices into scan data of the individual slices. Reference scan data must be determined for all the slices involved. This typically takes place in the context of a reference scan, performed in addition to the diagnostic scan, which measures the reference scan data individually for each required slice. Methods are also known in which, for determining the reference scan data, a separate reference scan is no longer necessary. Instead, for example, the reference scan is interlaced with the scanning of the slice multiplexing scan data, as described e.g. in United States Patent Application Publication No. 2018/0074147, or the reference data is determined from the slice multiplexing scan data, as described in the subsequently published EP17174507.
In order to be able to separate the resultant signals of the different slices, for example, a different phase (phase amount) is applied to each of the individual RF pulses (possibly before the multiplexing), for example by adding a different phase, in each case, which increases linearly (e.g. with the k-space coordinates in the phase encoding direction (ky)). In this way, each slice can be impressed with a different phase gradient, so that the slices are displaced against one another in the image space. This displacement is controlled by the so-called “field of view (FOV) shift factor” or “interslice FOV shift”. How an optimal FOV shift factor can be determined is described, for example, in the subsequently published DE102016218955.
In the CAIPIRINHA methods described in the aforementioned articles by Breuer et al. and Setsompop et al., by switching additional gradient blips or by additional modulation of the phases of the RF pulses of the multi-band RF pulses between the simultaneously excited slices, alternating further phase shifts are applied which generate displacements in the image space in the slice direction (“interslice FoV shifts”). These additional displacements in the image space improve the quality of the separation of the signals of the slices, in particular if the coil sensitivities have such slight differences in the sensitivity profiles of the individual coils used, that they are not sufficient for a reliable separation of the slices. Thus artifacts in the image data finally reconstructed from the recorded scan data are lessened.
The method that is most used to generate echo signals following an excitation of the nuclear spin is the so-called spin-echo method. In the simplest case, through radiation of at least one RF refocusing pulse following the radiation of the RF excitation pulse, the transverse magnetization is, so to speak, “turned” so that the dephazed magnetization is rephazed again and thus, following a time TE denoted as the echo time following the RF excitation pulse, a so-called spin echo SE is generated.
The excitation and scanning of the echo signals generated are repeated following a repetition time TR (e.g. by switching different gradients for position encoding) until the desired number of echo signals has been scanned and stored in k-space in order to image the examination object.
Among the SE sequences, in particular the TSE (“turbo spin echo”) sequences which are also known by the names FSE (“fast spin echo”) or RARE (“Rapid Acquisition with Refocused Echoes”) sequences, are widely used in clinical application. The advantage of the TSE sequences over the “simple” SE sequence is that following an RF excitation pulse, a plurality of refocusing pulses are switched and that thereby, a plurality of spin echo signals SE are generated following an excitation (multi-echo sequence). By this technique, the data recording is accelerated since fewer repetitions of the sequence with different position encoding are necessary to measure all the desired data. The scan time for the whole of the k-space in TSE sequences is thus reduced according to the number of the echo signals refocused and recorded following an excitation, the so-called “turbofactor” as compared with conventional SE methods.
By contrast, nuclear spin stimulated by an RF excitation pulse can be manipulated by switching dephazing and rephazing gradients so that the signal decays faster than is due to the T2* decay inherent to the scanned tissue, but after a particular time, the echo time TE, following the RF excitation pulse, a so-called measurable gradient echo forms. Such sequences are typically denoted GRE sequences. GRE sequences also include variants which generate a plurality of (gradient) echo signals following an excitation and therefore belong to the multi-echo sequences. Prominent variants are EPI (“echo planar imaging”) methods in which an oscillating readout gradient is used in which each change of the polarization direction of the gradient refocuses the transverse magnetization as far as the T2* decay allows, and thereby generates a gradient echo.
In the clinical use of magnetic resonance techniques, alongside the most often generated (contrast-weighted) MR images which represent the examination object in such a manner that anatomical information is presented very visibly, parameter maps that represent the local distribution of particular parameters of the examination object being imaged are also gaining importance. Possible parameters are, for example, relaxation parameters which reproduce the decay of the magnetization (in particular the longitudinal relaxation (T1-decay), the transverse relaxation (T2-decay) and the effective transverse relaxation (T2*-decay)), but also the proton density p or perhaps diffusion parameters.
In order to be able to create such parameter maps that represent the distribution of the proton density ρ or of the T1-decay constants or the T2-decay constants or the T2* decay constants, echo signals must be recorded at different echo times or inversion times (depending on the decay law; see next section), it is therefore useful to record this plurality of echo signals with a multi-echo sequence. From the scan data recorded at a common echo time or inversion time, an image data set can be created in each case. By fitting, such as pixel-by-pixel or voxel-by-voxel, the signal values obtained of the different image data sets to the respective decay laws (1) to (3), the local values of the parameters can be determined:
(1) T1-value and proton density ρ from:S=|ρ(1−2 exp(−TI/T1))|, with inversion time TI. 
(2) T2-value and proton density ρ from:S=ρ exp(−TE/T2), with echo time TE. 
(3) T2*-value and proton density ρ from:S=ρ exp(−TE/T2*), with echo time TE. 
A method for obtaining proton density (ρ)-maps, T1-maps, T2-maps and T2*-maps using a single slice EPI sequence is described by Poupon et al. in “Real-time EPI, T1, T2 and T2* mapping at 3T”, Proc. Intl. Soc. Mag. Reson. Med. 18, p. 4983, 2010, wherein EPI measurements with respective different inversion times were carried out successively in order to determine proton density maps and T1-maps (according to decay law (1)), spin echo EPI measurements, each with different echo times, to determine proton density maps and T2-maps (according to decay law (2)), and EPI scans with different echo times, for the determination of proton density maps and T2*-maps (according to decay law (3)), in each case of one slice, whereby the overall scan time was approximately 12 minutes.
For T1-maps, a preparation of the echo signals generated with corresponding RF inversion pulses is necessary. For other parameters, a preparation of the echo signals generated can also be necessary, i.e. the spins in or around a slice to be scanned are prepared in a desired manner before the excitation for generating the echo signals to be recorded with the aid of radiated RF pulses and/or switched gradients. This is the case, for example, for diffusion parameters.
In routine clinical practice, diffusion-weighted magnetic resonance (MR) images can provide important diagnostic information, for example in the case of diagnosing stroke and tumors. In diffusion-weighted imaging (DWI), for preparation, diffusion gradients are switched in particular directions, the diffusion of water molecules weakening the measured magnetic resonance signal along the diffusion gradients that are applied. Therefore, in regions with lower diffusion, a weaker signal attenuation takes place, so that these regions are imaged with a greater image intensity in an imaging magnetic resonance tomography (MRT) scan. The strength of the diffusion weighting is correlated to the strength of the applied diffusion gradients. The diffusion weighting can be characterized with the so-called b-value, which is a function of gradient parameters, for example, the gradient strength, duration or separation between the applied diffusion gradients. Due to the speed of these sequences, the recording of the resultant magnetic resonance signals usually takes place with a multi-echo sequence, for example, EPI.
During diffusion imaging, typically a number of images with different diffusion directions and weightings (characterized by the b-value) are recorded and combined with one another in order, for example, to calculate diffusion parameter maps, in particular of the diffusion parameters “Apparent Diffusion Coefficient” (ADC) and/or “Fractional Anisotropy” (FA). In the review article “Technical aspects of MR diffusion imaging of the body”, European Journal of Radiology 76, pp. 314-322, 2010, Dietrich et al. provide an overview of known DWI methods.
As stated above, slice multiplexing methods enable the scanning of echo signals from two or more slices simultaneously within one scan data recording. By this technique, in multi-echo methods in which following a first excitation of spin, through radiation of RF pulses and/or switching of gradients, a number of echo signals are generated and recorded. Compared with single slice scans, the minimum repetition time TR which is required for the recording of all the desired echo signals following a first excitation can be shortened, since fewer echo signals have to be created if they are recorded from a number of slices simultaneously than if only scan data of a single slice is contained in each recorded echo signal. For example, if scan data are measured from n slices simultaneously, the time required for recording the desired echo signals of all the slices is reduced by the factor n.
However, in many MR applications, the minimum repetition time TR is not determined by the time needed to record all the desired echo signals following an excitation, but rather is dependent on the respective desired image contrast. If a contrast weighted according to the T1, T2 or T2* decay constants of the respectively imaged tissue is to be achieved, the repetition time TR must be sufficient to be able to image the desired decay processes, since otherwise the desired image contrast is not determinable in sufficient quality from the scan data. Therefore, the known slice multiplexing methods produce no time advantage in multi-echo methods which are to generate the MR images with a particular contrast.
In United States Patent Application Publication Nos. 2017/0108567, 2017/0315202, 2018/0024214 and 2018/00316659, methods are described as to how, by slice multiplexing methods, scan data can be recorded, from which MR images with different contrast weightings can be generated. A further such method is described in the article “Simultaneous Multi-Contrast Imaging with ReadoutSegmented EPI”, Proc. Intl. Soc. Mag. Reson. Med. 25, p. 520, 2017 by Breutigam et al.
With that method, however, it is still not possible to record scan data in the shortest possible time from which different parameter maps can be determined.